Sheath flow is a widely used technique for a variety of applications, including but not limited to particle counting, flow cytometry, waveguiding, and fluid control. Sheath flow involves surrounding a central flow stream (the core) with a surrounding stream (the sheath). In particle counting and flow cytometry applications, the sheath prevents particles in the core from coming into contact with the walls of the channel, thus preventing adhesion and clogging. The sheath also serves to focus the particles or molecules into the center of the channel, allowing for easy counting or measurement through optical or other means. Sheath flow is normally laminar flow that substantially avoids mixing between the core stream and the sheath stream. Sheath flow can also be used with fluids of different refractive index to create a waveguide in the core or sheath stream in order to measure transfer of analytes from one stream to the other or to control the rate of interaction between molecules in one or both streams for carefully controlled chemistry or analysis.
Previous designs have created sheath flow through an annular arrangement. A small nozzle was positioned inside a larger tube. The core solution was pumped through the nozzle and the sheath solution was pumped through the larger tube. This configuration required careful alignment of the two tubes and did not easily lend itself to miniaturization. Since the diameter of the nozzle was fixed, the relative sizes of the core stream and sheath solution were relatively constant within a set range.
Other devices provide sheath flow on a chip, but the flow typically operates only in two dimensions. The core stream in these devices is bordered on either side by the sheath streams, however the core is not sheathed top and bottom. The complexity of the support plumbing for these devices is increased, as the number of flow streams is increased from two to three as compared to the annular arrangement design. It is possible to sheath the stream on the top and bottom of the core stream in these systems by adding two additional inlet ports in the top and bottom of the channel. However, this greatly increases the manufacturing complexity of the device. Micromachining technologies are inherently two-dimensional. Three-dimensional channel paths can be created by stacking several two dimensional designs on top of one another, but this adds to the complexity and difficulty of the manufacturing process. Creating a fully sheathed flow in this way could require at least several individual levels, which must be independently produced and then carefully aligned. In addition, use of the device could require multiple pumps to provide solutions to all the inlets.
Tissue Engineering
Studying the growth and differentiation of cells in culture in the presence of various nutrients and growth factors has provided physiologically relevant information about how the corresponding cells function in vivo. Classic tissue engineering mixes cell types with different growth factors and provides minimal control over morphology and microanatomy [Langer, R., Vacanti, J. P. 1993 “Tissue Engineering.” Science. 260(5110):920-6; Levenberg, S., Rouwkema, J., Macdonald, M., Garfein, E. S., Kohane, D. S., Darland, D. C., et al. 2005 “Engineering vascularized skeletal muscle tissue.” Nat Biotechnol. 23(7):879-84.]. It is becoming increasingly clear that cell differentiation and function are impacted by fluid flow, proximity of other cell types, and substrate geometry. The realization of the importance of such factors has motivated the microfluidics and tissue engineering communities to create “tissue-on-chip” or “organ-on-chip” model systems that can introduce methods for controlling such variables and provide more complex in vitro systems for the study of normal differentiation and pathogenesis or drug metabolism and transport (Wong, K. H. K., Chan, J. M., Kamm, R. D., Tien, J. 2012 Microfluidic Models of Vascular Functions. Annu Rev Biomed Eng. 14:205-230; Van Der Meer, A. D., Van Den Berg, A. 2012 “Organs-on-chips: breaking the in vitro impasse.” Integr Biol-UK 4(5):461-470; Shuler, M. L. 2012 “Modeling Life.” Ann. Biomed. Eng. 40(7):1399-1407).
In general, these tissue-on-chip models are configured with one of two types of architectures. For decades, investigators have been configuring substrates to have a specific geometry that will impact cell differentiation, most notably defining surface topography or creating channels to direct cell growth. In both cases, the cells are introduced after the substrate is configured, and cells adhere to defined portions of the substrate. Surface patterning of endothelial or progenitor cells has been used to engineer blood vessels [Niklason, L. E., Gao, J., Abbott, W. M., Hirschi, K. K., Houser, S., Marini, R., et al. 1999 “Functional arteries grown in vitro.” Science. 284(5413):489-93; Kaushal, S., Amiel, G. E., Guleserian, K. J., Shapira, O. M., Perry, T., Sutherland, F. W., et al. 2001 “Functional small-diameter neovessels created using endothelial progenitor cells expanded ex vivo.” Nat Med. 7(9):1035-40.]. Chip-based approaches to generating engineered blood vessels pattern layers of cells onto tubular or rectangular microchannels to emulate blood vessel geometry; however, these methods do not produce a free-standing engineered blood vessel. The microchannel-attached engineered blood vessel does not allow for superfusion along the blood vessel wall or branching from the main vessel. [Nichol, J. W., Koshy, S. T., Bae, H., Hwang, C. M., Yamanlar, S., Khademhosseini, A. 2010 “Cell-laden microengineered gelatin methacrylate hydrogels.” Biomaterials. 31(21):5536-44. Chau, L. T., Rolfe, B. E., Cooper-White, J. J. 2011 “A microdevice for the creation of patent, three-dimensional endothelial cell-based microcirculatory networks.” Biomicrofluidics. 5(3); Liu, Y. X., Markov, D. A., Wikswo, J. P., Mccawley, L. J. 2011 “Microfabricated scaffold-guided endothelial morphogenesis in three-dimensional culture.” Biomed Microdevices. 13(5):837-46.] More recently, porous membranes have been suspended across a microfluidic well or channel and cells grown on one or both surfaces of the porous membrane. The cells generally form monolayers while air or liquids are flowed above and/or below the membrane (for example, U.S. Patent Application Publication No. 2011/0250585A1 and Huh, D, Ingber et al., “Reconstituting organ-level lung functions on a chip,” Science, 2010, 328, 1662.). Both of these approaches limit the tissue-on-chip construct to planar configurations for the resulting cell organizations. The depth of the cell layers in planar tissue models is constrained by the need to transport nutrients and growth factors from the fluid in the microfluidic channel, preventing the formation of thick tissues.
In nature, the transport of nutrients, growth factors, protective molecules, and waste are provided by the vasculature and other types of ducts. Cells including both those that provide immunity and oxygenation and those that cause infection, autoimmune disease and cancer are also transported through the vasculature. The need for tubular structures to incorporate into tissue-on-chip models is recognized (e.g. Wong et al., ibid.). Ideally, such vasculature would be round as in nature, flexible to accommodate complex organ geometries, and composed of a biocompatible or biodegradable material that can be remodeled by the incorporated cells (not polydimethylsiloxane). The roadblock to vascularized tissue models is twofold: (1) accurate fabrication of engineered blood vessels and (2) integration of engineered blood vessels into on-chip models.
In the present state of the art, elongated blood vessel structures require either cells grown in channels [Franco, C., Gerhardt, H. 2012 “Tissue Engineering: Blood vessels on a chip.” Nature. 488(7412):465-6], cells seeded onto preformed scaffolds, or cultures using blood vessels excised from animals [Quint, C., Kondo, Y., Manson, R. J., Lawson, J. H., Dardik, A., Niklason, L. E. 2011 “Decellularized tissue-engineered blood vessel as an arterial conduit.” Proc. Nat'l Acad. Sci. USA. 108(22):9214-9]. The first methods do not provide stand-alone blood vessels and culturing excised vessels is not a viable source for producing cost and time-effective human tissue models. Production of collagen-based blood vessel scaffolds for subsequent cell incorporation is already being used clinically, where it is desirable for the patient to provide his own cells. However, these scaffolds are large and replace veins or arteries rather than capillaries. The scaffold structures are generally millimeters in diameter, and nutrient availability in vitro might limit cells to growth near the surface of the scaffold, making them less appropriate for tissue-on-chip models. Development of methods for producing blood vessels to operate and supply tissue models would provide a much more accurate model for blood delivery to tissue and provide for growth of on-chip tissues in three dimensions.
The range of applications for the nano- and microfabrication of cell-laden and/or coated tubular constructs are not limited to vasculature. Directed nerve growth has been shown whereby neurons are seeded into scaffolds composed of poly L-lactic acid nanofibers. [F. Yang, R. Murugan, S. Ramakrishna, X. Wang, Y.-X. Ma, S. Wang, “Fabrication of nano-structured porous PLLA scaffold intended for nerve tissue engineering,” Biomaterials 25 (2004) 1891-1900]. The scaffolds initiated axonal guidance, which was postulated as a first step in bridging the gaps between the proximal and distal nerves that do not close during healing. An inkjet printing station for neuroregenerative tissue engineering was presented by Silva, D. S., D. B. Wallace, et al. (2007) [IEEE Dallas Engineering in Medicine and Biology Workshop: 71-73]. The resulting tubes were seeded with neurons and served as scaffolds that resulted in significant outgrowth.
Ductal tissues play essential roles in human physiology by providing conduits that facilitate the transfer of fluids such as seminal fluid, bile, and milk. These conduits serve as the interfaces between exocrine glands and distal regions both internal and external. Ducts differ from blood vessels in that they are often not simply passive conduits, but actively participate in the secretions that facilitate the expelling of waste or transfer of reproductive materials. As an example, a multilayer microfluidic device that was used to culture and analyze renal tubule cells [Jang, K.-J., and Suh, K.-Y. “A multi-layer microfluidic device for efficient culture and analysis of tubular cells.” Lab on a Chip (2010) 10, 36-42]. A fibronectin-coated polyester membrane was inserted between the layers of a polydimethylsiloxane (PDMS) microfluidic channel, and primary kidney cells were introduced on one side. During culture, the membrane was subjected to continuous shear stress of 1 dyn/cm2 for 5 h. The cells formed a layer on the polyester membrane and developed markers typical of renal tubule cells.
In view of the above, a need exists for the engineering of blood vessels, tissue ducts, and the like having physiologically-appropriate shapes, dimensions, and properties. Most prior methods for making polymer fibers involve conditions that are not compatible with living cells, in that they typically involve elevated temperature, high sheer, organic solvents, or combinations of these. Techniques described herein provide a biocompatible method for making shaped polymer fibers using hydrodynamic focusing.